This application is a divisional of U.S. application Ser. No. 10/244,306, filed on Sep. 16, 2002, the disclosure of which is incorporated herein by reference.
This invention relates to a prosthetic intervertebral disc nucleus. More particularly, it relates to an artificial disc nucleus made of a hydrogel material having a radiovisible material therein.
The intervertebral disc is a complex joint anatomically and functionally. It is composed of three component structures; the nucleus pulposus (the nucleus), the annulus fibrosus (the annulus) and the vertebral end-plates. The biochemical composition and anatomical arrangements within these component structures are related to the biomechanical function of the disc.
The nucleus occupies about 25-40% of the total disc cross-sectional area. It is primarily composed of mucoid material containing mainly proteoglycans with a small amount of collagen. The proteoglycans consist of a protein core with chains of negatively charges keratin sulphate and chondroitin sulphate covalently attached thereto. Due to these constituents, the nucleus is a loose hydrogel which usually contains about 70-90% water by weight. Although the nucleus plays an important role in the biomechanical function of the disc, the mechanical properties of the disc are not well known, largely because of the loose hydrogel nature of the nucleus.
As the nucleus is surrounded by the annulus and vertebral end-plates and the negatively charged sulphate groups are immobilized due to the attachment of these groups to the polymer matrix, the matrix has a higher concentration of counter ions than its surroundings. This ion concentration results in a higher osmotic pressure than the annulus e.g., ranging from about 0.1 to about 0.3 MPa. As a result of the high fixed charge density of the proteoglycan the matrix exerts an osmotic swelling pressure which can support an applied load in much the same way as air pressure in a tire supports the weight of a car.
It is the osmotic swelling pressure and hydrophilicity of the nucleus matrix that offers the nucleus the capability of imbibing fluid until it is balanced with the internal resistance stresses, due to the tensile forces of the collagen network, and the external stresses due to the loads that are applied by muscle and ligament tension. The swelling pressure (Ps) of the nucleus is directly dependent on the concentration and fixed charge densities of proteoglycan, i.e., the higher the concentration and fixed charge densities of proteoglycan the higher will be the swelling pressure of the nucleus. The external pressure changes with posture. When the human body is supine the compressive load on the third lumbar disc is 300 newton (N) which rises to 700 N when an upright stance is assumed. The compressive load increases, yet again, to 1200 N when the body is bent forward by only 20° C. When the external pressure (Pa) increases the previous balance, i.e. Ps=Pa, is upset. To reach a new balance, the swelling pressure has to increase. This increase is achieved by increasing the proteoglycan concentration in the nucleus which is achieved by reducing the fluid in the nucleus. That is why discs lose about 10% of their height, as a result of creep, during the daytime. When the external load is released i.e., Ps is greater than Pa, the nucleus will imbibe fluid from its surroundings in order to reach the new equilibrium value. It is this property of the nucleus that is mainly responsible for the compressive properties of the disc.
The annulus forms the outer limiting boundary of the disc. It is composed of highly structured collagen fibers embedded in an amorphous base substance which is also composed of water and proteoglycans. The amount of proteoglycans is lower in the annulus than in the nucleus. The collagen fibers of the annulus are arranged in concentric laminated bands or lamella, (about 8-12 layers thick) with a thicker anterior wall and thinner posterior wall. In each lamella, the fibers are parallel and attached to the superior and inferior vertebral bodies at an angle of about 30° form the horizontal plane of the disc in both directions. This design particularly resists twisting because the half of the fibers cocked in one direction will tighten as the vertebrae rotate relative to each other in the other direction. The composition of the annulus along the radial axis is not uniform. There is a steady increase in the proportion of collagen from the inner to the outer sections of the annulus. This difference in composition may reflect the need of the inner and outer regions of the annulus to blend into very different tissues while maintaining the strength of the structure. Only the inner lamellae are anchored to the end-plates forming an enclosed vessel for the nucleus. The collagen network of the annulus restrains the tendency of the nucleus gel to absorb water from surrounding tissues and swell. Thus, the collagen fibers in the annulus are always in tension, and the nucleus gel is always in compression.
The two vertebral end-plates are composed of hyaline cartilage, which is a clear, “glassy” tissue, that separates the disc from the adjacent vertebral bodies. This layer acts as a transitional zone between the hard, bony vertebral bodies and the soft disc. Because the intervertebral disc is avascular, most nutrients that the disc needs for metabolism are transported to the disc by diffusion through the end-plate area.
The intervertebral joint exhibits both elastic and viscous behavior. Hence, during the application of a load to the disc there will be an immediate “distortion” or “deformation” of the disc, often referred to as “instantaneous deformation.” It has been reported that the major pathway by which water is lost, from the disc during compression, is through the cartilage end-plates. Since the water permeability of the end-plates is in the range of about 0.20 to about 0.85×10−17m4N−1 sec−1 it is reasonable to assume that under loading, the initial volume of the disc is constant while the load is applied. Because the natural nucleus of the disc is in the form of a loose hydrogel, i.e., a hydrophilic polymeric material which is insoluble in water, it can be deformed easily, the extent of deformation of the disc being largely dependent on the extensibility of the annulus. It is generally believed that hydrostatic behavior of the nucleus plays an important role in the normal static and dynamic load-sharing capability of the disc and the restoring force of the stretched fibers of the annulus balances the effects of the nucleus swelling pressure. Without the constraint by the annulus, annular bulging of the nucleus would increase considerably. If the load is maintained at a constant level, a gradual change in joint height, commonly referred to as “creep” will occur as a function of time. Eventually, the creep will stabilized and the joint is said to be in “equilibrium.” When the load is removed the joint will gradually “recover” to its original height before loading. The creep and relation rates depend on the amount of load applied, the permeability of the end-plates and the water binding capability of the nucleus hydrogel. Creep and relaxation are essential processes in pumping fluid in and out of the disc.
Degeneration of the intervertebral disc is believed to be a common cause of final pathological changes and back pain. As the intervertebral disc ages it undergoes degeneration. The changes that occur are such that, in many respects, the composition of the nucleus seems to approach that of the inner annulus. Intervertebral disc degeneration is, at least in part, the consequence of compositional changes in the nucleus. It has been found that both the molecular weight and the amount of proteoglycans in the nucleus decrease with age, especially in degenerated discs, and the ratio of keratin sulphate to chondroitin sulphate in the nucleus increases. This increase in the ratio of keratin sulphate to chondroitin sulphate and decrease in proteoglycan content decreases the fixed charge density of the nucleus from about 0.28 meq/ml to about 0.18-0.20 meq/ml. These changes cause the nucleus to lose part of its water binding capability which decreases the maximum swelling pressure it can exert. As a result, the maximum water content drops from over about 85%, in preadolescence, to about 70-75% in middle age. The glycosaminoglycan content of prolapsed discs has been found to be lower, and the collagen content higher, than that of normal discs of a comparable age. Discs L-4-L-5 and L-5-S-1 are usually the most degenerated discs.
It is known that although the nucleus only occupies about one third of the total disc area, it takes about 70% of the total loading in a normal disc. Thus, it has been found that the compressive load on the nuclei of moderately degenerated discs is about 30% lower than in comparable normal discs but the compressive load on the annulus increases by 100% in the degenerated discs. This load change is primarily caused by the structural changes in the disc as discussed above. The excess load on the annulus, of the degenerated disc, causes reduction of the disc height and excessive movement of the spinal segments. The flexibility of the disc produces excessive movement of the collagenous fibers which in turn, injures the fiber attachments and causes delamination of the well organized fibers of the annulus ring. The delamination annulus can be further weakened by stress on the annulus and in severe cases this stress will cause tearing of the annulus. This whole process is very similar to driving on a flat tire, where the reinforcement layer will eventually delaminate. Because the thickness of the annulus is not uniform, with the posterior portions being thinner than the anterior portions, delamination and lesions usually occur in the posterior area first.
The spinal disc may also be displaced or damaged due to trauma or diseases. In these cases, and in the case of disc degeneration, the nucleus may herniate and/or protrude into the vertebral canal or intervertebral foramen, in which case it is known as a herniated or “slipped” disc. This disc may in turn press upon the spinal nerve that exits the vertebral canal through the partially obstructed foramen, causing pain or paralysis in the area of its distribution. The most frequent site of occurrence of a herniated disc is in the lower lumbar region. A disc herniation in this area often involves the inferior extremities by compressing the sciatic nerve.
There are basically three types of treatment currently being used for treating low back pain caused by injured or degenerated discs: conservative care, discectomy and fusion. Each of these treatments has its advantages and limitations. The vast majority of patients with low back pain, especially those with first time episodes of low back pain, will get better with conservative treatment. However, it is not necessarily true that conservative care is the most efficient and economical way to solve the low back pain problem.
Discectomy usually provides excellent short term results in relieving the clinical symptoms, by removing the herniated disc material, usually the nucleus, which causes the low back pain either by compressing the spinal nerve or by chemical irritation. Clearly, a discectomy is not desirable from a biomechanical point of view. In a healthy disc, the nucleus takes the most compressional load and in a degenerated disc this load is primarily distributed onto the annulus ring which, as described above, causes tearing and delamination of the annulus. Removal of the nucleus in a discectomy actually causes distribution the compressive load onto the annulus ring thereby narrowing the disc spaces. It has been reported that a long-term disc height decrease might be expected to cause irreversible osteoarthritis-like changed in the facet joint. That is why discectomy yields poor long term benefits and results in a high incidence of reherniation.
Fusion generally does a good job in eliminating symptoms and stabilizing the joint. However, because the motion of the fused segment is restricted, the range of motion of the adjoining vertebral discs is increased possibly enhancing their degenerative processes.
Because of these disadvantages, it is desirable to use a prosthetic joint device which not only is able to replace the injured or degenerated intervertebral disc, but also can mimic the physiological and the biomechanical function of the replaced disc and prevent further degeneration of the surrounding tissue.
Artificial discs are well known in the prior art. U.S. Pat. No. 3,867,728, to Stubstad et al., relates to a device which replaces the entire disc. This device is made by laminating vertical, horizontal or axial sheets of elastic polymer. U.S. Pat. No. 4,309,777, to Patil, relates to a prosthetic utilizing metal springs and cups. A spinal implant comprising a rigid solid body having a porous coating on part of its surface is shown in Kenna's U.S. Pat. No. 4,714,469. An intervertebral disc prosthetic consisting of a pair of rigid plugs to replace the degenerated disc is referred by Kuntz, U.S. Pat. No. 4,349,921. U.S. Pat. Nos. 4,772,287 and 4,904,260 to Ray et al., teach the use of a pair of cylindrical prosthetic intervertebral disc capsules with or without therapeutical agents. U.S. Pat. No. 4,911,718 to Lee et al., relates to an elastomeric disc spacer comprising three different parts; nucleus, annulus and end-plates, of different materials. At the present time, none of these inventions has become a product in the spinal care market. Bao et al., in U.S. Pat. Nos. 5,047,055 and 5,192,326 (assigned to the assignee of this invention and incorporated herein by reference) describe artificial nuclei comprising hydrogels in the form of large pieces shaped when fully hydrated, to generally conform to the disc cavity or hydrogel beads within a porous envelope, respectively. The hydrogels have an equilibrium water content (EWC) of at least about 30% and a compressive strength of at least about 1 meganewtons per square meter (1 MNm−2) when subjected to the constraints of the annulus and end plates of the disc. Preferably, the compressive strength of the nucleus is about 4 MNm−2 or even higher.
The primary disadvantage of the invention of Substad et al., Patil, Kenna and Lee et al., is that use of their prosthesis requires complete replacement of the natural disc which involves numerous surgical difficulties. Secondly, the intervertebral disc is a complex joint, anatomically and functionally, comprising the aforementioned three component structures, each of which has its own unique structural characteristics. Designing and fabricating such a complicated prosthesis from acceptable materials, which will mimic the function of the natural disc, is very difficult. A further problem is the difficulty of preventing the prosthesis from dislodging. Fourthly, even for prostheses which are only intended for replacing the nucleus, a major obstacle has been to find a material which is similar to the natural and is also able to restore the normal function of the nucleus. Hydrophobic elastomers and thermoplastic polymers are not desirable for use in the prosthetic nuclei due to their significant inherent differences from the natural nucleus e.g., lack of hydrophilicty, in the elastomers, and lack of flexibility in their thermoplasts.
These problems are not solved by Kuntz, who uses elastic rubber plugs, or by Froning and Ray et al., who use bladders, or capsules, respectively, which are filled with a fluid or thixotropic gel. According to the Ray and Froning patents, liquid was used to fill the capsules and bladders, respectively, thereby requiring that their membranes be completely sealed to prevent fluid leakage. As a consequence, those devices cannot completely restore the function of the nucleus which allows body fluid to diffuse in and out during cyclic loading thereby providing the nutrients the disc needs.
The Bao et al., prosthetic lumbar disc nuclei are made from hydrogels. Hydrogels have been used in biomedical applications, such as contact lenses. Among the advantages of hydrogels is that they are more biocompatible than hydrophobic elastomers and metals. This biocompatibility is largely due to the unique characteristics of hydrogels in that they are soft and contain water like the surround tissues and have relatively low frictional coefficients with respect to the surrounding tissues. The biocompatibility of hydrogels results in prosthetic nuclei which are more easily tolerated in the body. Furthermore, hydrophobic elastomeric and metallic gels will not permit diffusion of aqueous compositions, and their solutes, therethrough.
An additional advantage of some hydrogels is their good mechanical strength which permits them to withstand the load on the disc and restore the normal space between the vertebral bodies. The aforementioned nuclei of Bao et al. have high mechanical strength and are able to withstand the body loads and assist in the healing of the defective annuli.
Other advantages of the hydrogels, used in Bao et al. nuclei, are their excellent viscoelastic properties and shape memory. Hydrogels contain a large amount of water which acts as a plasticizer. Part of the water is available as free water which has more freedom to leave the hydrogel when the hydrogel is partially dehydrated under mechanical pressure. This characteristic of the hydrogels enables them to creep, in the same way as the natural nucleus, under compression and to withstand cyclic loading for long periods without any significant degradation or loss of their elasticity. This is because water in the hydrogel behaves like a cushion whereby the polymeric network of a hydrogel with a high equilibrium water content (EWC) is less susceptible to damage under mechanical load.
Another advantage of hydrogels is their permeability to water and water-soluble substances, such as nutrients, metabolites and the like. It is know that body fluid diffusion, under cycle loading, is the major source of nutrients to the natural disc. If the route of this nutrient diffusion is blocked, e.g., by a water-impermeable nucleus, further deterioration of the disc will ensure.
Hydrogels can be dehydrated and the resultant xerogels hydrated again without changing the properties of the hydrogels. When a hydrogel is dehydrated, its volume decreases, thereby facilitating implantation of the prosthetic nucleus into the nuclear cavity in the disc. The implanted prosthetic nucleus will then swell, in the body, by absorption of body fluid up to its EWC. The EWC of the hydrogel depends on the compressive load applied thereto. Thus, the EWC of a specific hydrogel in an open container will differ from the EWC of the same hydrogel in a closed vessel such as an intervertebral disc. The EWC values, referred to below, are for hydrogels subjected to compressive loads under the conditions found in an intervertebral disc. The expansion factor of a dehydrated hydrogel, in turn, is dependent on its EWC. Thus, it may vary from 1.19 for a hydrogel of 38% EWC to 1.73 for a hydrogel of 80% EWC. For an 80% EWC hydrogel, the volume of the dehydrated prosthetic nucleus is usually about 20% of that of the hydrated one. The ability to be dehydrated and then return to its original shape upon hydration, up to its EWC, makes it possible to implant the device posterior-laterally during surgery, thereby reducing the complexity and risk of intraspinal surgery as traditionally used. The danger of perforation of the nerve, dural sac, arteries and other organs is also reduced. In addition, the incision area on the annulus can be reduced, thereby helping to heal the annulus and prevent the reherniation of the disc. Hydrogels are also useful for drug delivery into the disc due to their capability for controlled release of drugs. Various therapeutic agents, such as growth factors, long term analgesics and anti-inflammatory agents can attach to the prosthetic nucleus and be released in a controllable rate after implantation of the nucleus in the disc.
Furthermore, dimensional integrity can be maintained with hydrogels having a water content of up to about 90%. This dimensional integrity, if the nucleus is properly designed will aid in distributing the vertebral load to a larger area on the annulus ring and prevent the prosthetic nucleus from bulging and herniating.
However, it is normally difficult to implant a fully hydrated hydrogel prosthesis in the cavity, of a disc, through the small window provided in the disc, for removing the herniated nucleus, especially in a percutaneous surgery by virtue of their bulkiness in a fully in a fully hydrated state. Therefore, such prosthesis must be implanted, in the disc in relatively dehydrated states which requires long periods to achieve their EWCs due to their low surface areas. Other hydrogels, having high surface areas, do not completely conform to the shape of the nuclear cavity. Other polymers such as those disclosed in WO 97/268407 (PCT/US97 00457), the teachings of which are incorporated herein by reference, can also be used to fill the disc nucleus.
It is desirable to provide a hydrogel implant which is inherently radiopaque, i.e., radiovisible so that surgeons could view the placement of the implant in the cavity produced by the removal of a spinal nucleus. It is advantageous if the radiovisible material could be incorporated into the polymeric or hydrogel material making up the prosthetic nucleus implant. It is desirable to have a method of making the hydrogel or polymer radiopaque which would allow dimensional changes in the hydrogel implant during processing and after implantation without compromising the mechanical integrity of the implant.
Various methods are used to implant a hydrogel or other polymeric nucleus implant. Such a method is shown in U.S. Pat. No. 5,800,549, the teachings of which are incorporated herein by reference.